Embodiments of the invention relate generally to diagnostic imaging and, more particularly, to a modular multispot x-ray source for use in an imaging system.
Traditional x-ray imaging systems include an x-ray source and a detector array. X-rays are generated by the x-ray source, passed through and attenuated by an object, and are detected by the detector array. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the object. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis, which ultimately produces an image.
Generally, as in a CT application, the x-ray source and the detector array are mounted on a gantry and rotated about an imaging plane and around the object. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator adjacent the collimator for converting x-rays to light energy, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom. The X-ray detectors may also include a direct conversion device for discriminating the energy content of the x-ray beam. The outputs of the detector array are then transmitted to the data processing system for image reconstruction. Electrical signals generated by the detector array are conditioned to reconstruct an x-ray image of the object.
In CT imaging systems, the gantry rotates at various speeds in order to create a 360° image of the object. The gantry contains an x-ray source having an electron source or cathode assembly that generates electrons that are accelerated across a vacuum gap to a target or anode assembly via a high voltage potential. In releasing the electrons, a filament contained within the electron source is heated to incandescence by passing an electric current therethrough. The electrons are accelerated by the high voltage potential and impinge upon a target surface of the target at a focal spot. Upon impingement, the electrons are rapidly decelerated, and in the process, x-rays are generated therefrom.
The process of deceleration typically results in heating of the focal spot to very high temperatures. Thus, x-ray tubes include a rotating target or anode structure for the purpose of distributing heat generated at the focal spot. The target is typically rotated by an induction motor having a cylindrical rotor built into a cantilevered axle that supports a disc-shaped target and an iron stator structure with copper windings that surrounds an elongated neck of the x-ray tube. The rotor of the rotating target is driven by the stator. Because of the high temperatures generated when the electron beam strikes the target, the target is typically rotated at high rotational speed.
Newer generation x-ray tubes have increasing demands for providing higher peak power, thus generally higher average power as well. Higher peak power, though, would result in higher peak temperatures occurring in the target, particularly at the “track” or the point of impact on the target, unless the target design is altered. Because x-ray tubes are typically designed having peak temperatures at limits imposed by material capabilities and high voltage considerations, higher peak power typically calls for a re-design of the target. For a rotating target, the re-design may include higher rotation speed, larger track radius, or novel x-ray production means. These designs may reduce life and reliability of the rotating target. For stationary target sources, the re-design options are generally limited to material improvements or novel approaches to backscattered electron energy management.
Furthermore, newer generation CT systems have increased gantry speed requirements to better enable, for instance, cardiac imaging. Thus, systems have been designed having applications wherein the gantry is spun at or below 0.5 seconds rotational speed. Such applications may include yet faster gantry rotation, thereby increasing the g-load demands to, for instance, 0.2 second rotation, which represents a g-load well in excess of what can be withstood in many current CT systems.
Accordingly, to counter the need for high g-load capability x-ray sources, multispot systems have been designed having stationary imaging components therein. For instance, scanning electron beam (e-beam) x-ray sources include an electron gun positioned at a gantry center that emits an e-beam that is magnetically deflected toward a target. In such a system, the target typically forms a continuous ring surrounding a patient, and the e-beam is rapidly deflected to circumferential locations on the target and around the patient. The e-beam may be deflected in the z-direction as well. As such, multispot imaging may be performed very rapidly using stationary components. However, not only are such systems expensive, they may be prone to performance degradation as well. For instance, the continuous target may have thermal distortion that can degrade image quality through excessive focal spot motion.
Furthermore, other known systems having stationary components include a thin transmission-style target for x-ray generation. However, such a continuous target is likewise prone to thermal loading and distortion effects resulting, as well, in degraded image quality through excessive focal spot motion.
As such, modular multispot devices have been developed to reduce the thermal distortion effects resulting from large, continuous targets or anodes. In such a system, individual, modularized x-ray sources may be positioned within a gantry, each module having a plurality of individual or discrete focal spots that have reduced relative motion. As such, the overall system thermal distortion may be minimized and image quality may be improved. A modular design has the benefit of simplifying manufacturing and assembly procedures because the individual modules may be assembled and tested as sub-units before being installed into the overall system. Such a design further simplifies troubleshooting and repair of the system in the field, as a field engineer may be able to test and replace individual modules within the system. Thus, the need to return all of the sources or even the entire system back to a manufacturing site may be precluded, resulting less in system downtime, cost of repair, and frustration.
However, a multispot source typically results in the need to provide x-ray shielding of many spatially distributed focal spots. Adopting a traditional shielding approach would require covering the vacuum chamber containing the modules with lead or other high-density shielding material to eliminate the openings from which undesired x-rays could emanate. This presents at least two issues: first, the basic amount of shielding material would be large; and second, the amount of scattered radiation produced by objects inside the vacuum chamber makes the determination of the minimum thickness of shielding material required at all locations difficult.
Thus, not only is the basic amount of shielding material prohibitive, but because of the variation from system to system and the resulting uncertainty of sources of scattered radiation and to be conservative, designs typically include excess amounts of shielding. This results in increased system cost and an unnecessary amount of shielding mass being included in the system. As such, the desire for increased g-load capability may be limited due to the excess shielding required in a modular source design.
Furthermore, modular source designs typically include a pre-patient collimator to collimate scatter and off-focal radiation that may emit from the anodes. However, to collimate each spot within a multispot source, a separate collimator is provided for each spot, resulting in a series of individually constructed collimators. Further, in order to collimate in both the X and Z dimensions, respective collimating plates or elements must be provided in each orientation. Such a construction is complex and expensive to build, and cumbersome and difficult to operate.
Therefore, it would be desirable to design a collimator that collimates each spot in both X and Z dimensions for a modular multispot x-ray source.